According to the prior art, positron-emission topography detector rings are used in order to detect β+β− annihilation radiation. The rings are composed of scintillation crystals adjoined by sensors that are capable of detecting scintillation radiation. Typical detectors are SiPMs (silicon photomultipliers). The construction is such that the detector ring is generally circular and the object to be measured—such as a human or animal patient's body part, for example—is placed in the center of the detector ring (PET ring). β+β− annihilation radiation is produced through the use of radiodiagnostics, and that radiation is to be detected. The β+β− annihilation radiation, hereinafter called annihilation radiation, strikes scintillation crystals that are arranged in the manner of a ring around the object to be studied and produces the scintillation radiation. The scintillation radiation, in turn, is registered by the SiPM, which is located in the concentric arrangement behind the scintillation crystal in relation to the radiation source. However, the SiPMs can also be arranged on other sides of the scintillation crystal—in front of the scintillation crystal or to the side thereof, for example. The scintillation crystal is a three-dimensional body. In relation to an arrangement in which the object to be studied emits annihilation radiation from the center of the detector ring, the cross section upon which the annihilation radiation is incident on the scintillation crystal corresponds an xy axis. The depth of the scintillation crystal is referred to in this nomenclature as the z axis. In an idealized representation, an object to be studied or an emission source for radiating an energy of 511 keV, which ideally strikes the xy plane of the scintillation crystal perpendicularly and has a depth of penetration along the z axis of the scintillation crystal, is located in the center of the detector ring. The 511 keV annihilation radiation then triggers scintillation at a point of the scintillation crystal along the z axis that is registered by the sensor—an SiPM, for example—as a signal. An SiPM is even capable of detecting individual photons. When the minimum required light strikes the active sensor surface, the SiPM microcell experiences diode breakdown. This generates a current pulse, which can be measured at the output of the component. A so-called quench resistor prevents the cell from generating a critical current that is so high that the component is destroyed. The output current of an SiPM microcell is independent of the quantity of light that reached the sensor and started the breakdown process. An SiPM microcell is a binary sensor that detects whether or not light is incident. In order to obtain quantitative information about the incident light, an SiPM is composed of a plurality of microcells. A microcell consists of a photodiode and a quench resistor. The number of broken-down cells then provides information about the quantity of incident light.
A correlation exists between the sensitivity of the scintillation crystal and the length thereof along the z axis. The deeper the dimensioning of the scintillation crystal, the more sensitive it is, since the occurrence of a scintillation event becomes all the more likely. During the detection of the annihilation radiation, beams are emitted in two opposite directions from the point at which the annihilation radiation is emitted, so that the beams form a 180° angle. The line that is formed by these beams is referred to as the “line of response” (LOR). Accordingly, in the case of a ring-shaped detector, two beams strike scintillation crystals along the LOR that lie on opposing sides with respect to the ring-shaped arrangement, in the center of which the emission source is located.
For detectors with light detection by means of photodiodes in the form of SiPMs on only one side of the scintillation crystal, various established methods exist for determining the x and y position of an event. However, these do not include the z position, so the exact position in the scintillation crystal where the gamma photon was stopped on the z axis and converted into light is not determined. If the z position is not determined, parallax errors occur in the determination of the LOR that can be attributed to the so-called depth-of-interaction (DOI) problem. The DOI problem always arises when the point from which the emission of the annihilation radiation emanates does not lie exactly in the center of a ring-shaped detector. The farther the emission center for an LOR lies outside of the center of a PET ring, the greater the problem becomes. As a result, when a PET ring is designed, a compromise is made between an increase in sensitivity through longer scintillation crystals and a reduction in DOI errors through shorter scintillation crystals. In some areas of the PET application, the need exists to use PET rings (detector rings) that lie closely against the object of study. That is the case in medicine, in particular, if patients are to be examined simultaneously using an MRT method and a PET method. In that case, the PET ring must fit into the opening of the MRT scanner tube. One consequence of this is that the PET ring that is used must have a small diameter in order to fit into the opening of the MRT ring. However, if the dimensions of the PET ring are small, the problem arises that, while the object to be studied, such as a body part of a small animal or also of a human being, can be arranged so as to be centered, its dimensions will extend in relation to the diameter of the PET ring so as to reach far into the edge regions of the opening of the PET ring. This means, however, that points from which annihilation radiation emanates are also positioned so close to the PET ring that the DOI problem becomes substantial.
In previous years, the use of pixelated scintillation crystal blocks with smaller and smaller pixel sizes has resulted above all in a substantial improvement in the resolution in PET scanners for small animals. Pixeling is thus achieved on the xy plane, so that tubes of pixels that are aligned in the z direction form in the scintillation crystal. This was stimulated particularly by the need for higher and higher spatial resolution in PET scanners for small animals, since the object being examined is very small. Meanwhile, pixel size has already reached the sub-millimeter range. For this reason, two problems are arising more and more that must be resolved. Firstly, the pixelated crystal blocks are made of adhesive and reflector film, which is located between the individual scintillation crystals, in order to construct the pixelated block. The layer of adhesive and reflector film has an approximate thickness of 70 μm. Consequently, pixelated arrays with an especially small pixel pitch suffer increased loss of sensitivity. In the case of an array with crystal pixels measuring 0.8 cm×0.8 cm, such as those used in [1], for example, the ratio of adhesive and film to scintillation crystal decreases substantially, so that adhesive and film already make up a proportion of 29%. The proportion of scintillation crystal is logically reduced to 71%. No gamma quanta can be stopped and converted into light in the other 29% of the volume. If even smaller pixelated arrays measuring 0.5 cm×0.5 cm are used, for example, the proportion of crystal is even reduced to 59%. With pixelated arrays, an increase in resolution is therefore always associated with a loss of sensitivity. The second problem with pixelated scintillation crystal arrays is that the emitted light is concentrated on a smaller region of the SiPM detector surface. An SiPM consists of a plurality of microcells that act as binary elements. They identify whether or not light has been detected. If light has been detected, the microcell carries out a breakdown. The number of broken-down microcells indicates quantitatively how much light has reached the detector surface. If two or more light quanta trigger a microcell, the output signal remains the same. The more light is incident on an SiPM, the greater the likelihood that two or more light quanta will strike the same microcell of the SiPM. These additional light quanta then cannot be detected. Consequently, the likelihood of the saturation of a microcell is substantially greater if pixelated scintillation crystal arrays are used, since these concentrate the light more strongly on a small region of the sensor. Saturation effects also result in poorer energy resolution of the detectors.
As mentioned at the outset, detectors from the prior art use SiPM-based sensor technologies in order to enable magnetic resonance tomography (MRI) compatibility for use in MR/PET hybrid scanners. Another problem with hybrid scanners is that the space for PET detectors and associated electronics is limited by the tube diameter of the magnetic resonance tomograph (MRT). This applies particularly to ultra-high field tomographs. As a consequence of the narrower tube diameter, the PET scintillation crystals must be as short as possible. Shorter scintillation crystals also reduce the sensitivity. This also means that, due to the constraints of the tube diameter, the PET ring is located closer to the object of study. The closer the annihilations and hence the resulting LOR take place to the PET ring, the greater the parallax error. This is because the gamma quanta are no longer perpendicularly incident into the scintillation crystals if the annihilation occurs near the PET ring. With respect to PET ring design, this means that the parallax errors increase and become more pronounced when the PET ring is located close to the object to be studied, since annihilations can also take place near the PET ring in that case. Irrespective of limitations posed by hybrid devices, one also aims to design the PET rings so as to be as narrow as possible for the sake of greater sensitivity and lower costs.
Furthermore, it is known that many SiPM sensor concepts include the encoding of the output channels, since the power consumption of the PET ring is increased by an increase in output channels. However, this is limited by construction-related aspects. A simple calculation makes this clear. A PET ring with a diameter of 8 cm and a length of 10 cm results in a detector surface of 251 cm2. If a 1-to-1 coupling of scintillation crystals and SiPMs with a crystal pixel size of 0.8 mm is used, 39,270 readout channels are already required if each channel is to be read out individually.
In order to achieve greater spatial resolutions, current sensor designs consist of sensor chips with narrower pixel sizes. This results in a substantial increase in readout channels, which are limited by the power consumption, space, and data rates. Consequently, position-sensitive (PS) encoding methods have been developed in order to reduce the number of readout channels of a chip [1-3, 15]. The most current concept to be developed is called PS-SSPM [1] and is based on charge-sharing PS-SiPMs. Charge-sharing PS-SiPM microcells detect light like conventional SiPM microcells. However, this sensor design includes a resistor network that distributes the generated charge as a function of the position and the coding. The detector construction presented in [1] consists of a pixelated crystal array with a pitch of 0.8 mm.
This most up-to-date detector concept enables the advantage of a reduction in output channels through the channel coding to be achieved simultaneously with a high detector array resolution, which is achieved through the use of pixelated scintillation crystal arrays with a pitch of less than one millimeter. However, it does not include any DOI information detection.
One concept published in [4] proves that it is possible to construct a PET detector consisting of monolithic crystals and SiPMs. As mentioned previously, monolithic crystals solve the problem of losses of sensitivity due to the space requirement of reflector films and associated adhesives. Moreover, the production costs of monolithic crystals are lower as a result. The thickness that is used for the crystals is 2 mm. Parallax errors are thus prevented with the construction used in [4], but that comes at the price of the small extension of the scintillation crystal in the z direction. At the same time, the detection efficiency is low as a result of the low crystal height.
There are various possibilities for measuring DOI information and thus correcting parallax errors, which also detect light on another crystal side. Particularly for SiPMs of the prior art, this increases the costs immensely. One concept for DOI detection that detects light on only one crystal side and uses monolithic crystals is published in [5] and patented in [6]. It makes use of the known principle that the light distribution of the crystal is dependent on the DOI. The detector concept that is used is coupled with monolithic crystals to H8500 position-sensitive photomultipliers (PMTs) by Hamamatsu. Moreover, a resistor network is used that enables position encoding and thus also a reduction in output channels. The standard deviation of the light distribution is used in order to estimate the DOI. In order to calculate the standard deviation, the first- and second-order moment of the light distribution is needed. The first-order moment is already given by the linear coding of the output channels. A summation network has been developed and integrated into the resistor network in order to determine the second-order moment.
A summary overview of PET detectors with DOI detection is provided in [7]. Descriptions and results of PET- and MR/PET hybrid scanners for small animals that have been developed in recent years can be found in [8-11].
Detector concepts that are based on current SiPM-based technology and include position coding for channel reduction do not include any DOI detection. For this reason, PET rings that are constructed with these detectors include parallax errors in the reconstruction. What is more, most of them use pixelated crystal arrays. As described above, this results in a loss of sensitivity caused by the reflector film and the adhesive between the crystals of the array. Due to the lack of DOI information, the thickness of the crystals is limited. An increase in sensitivity by means of thicker crystals is accompanied by a loss in spatial resolution caused by the lack of DOI information. The DOI concepts for pixelated crystals cited in [7] cannot be used with crystals of arbitrarily small size and do not work for crystal arrays with crystal sizes of 0.8 mm or 0.5 mm. The main problem with the lack of DOI detection is that the PET ring size is limited, and a narrower ring would reduce the spatial resolution.
The detector described in [4] is embodied with monolithic crystals. A closely adjoining ring was designed in order to increase the sensitivity. At the same time, monolithic crystals were used. Due to the resulting short distance between the scintillation crystals and the object of study, the DOI problem is exacerbated. Consequently, the developers of the ring are limited to a 2 mm crystal thickness. This has the consequence that the sensitivity that is gained by the narrow ring and the use of monolithic crystals is lost again as a result of the small thickness of the scintillation crystals. That having been said, that work demonstrates that a high resolution is possible with monolithic crystals.
DOI positions with SiPM-based detectors can be determined by mounting sensors on two crystal faces. This requires twice the SiPM sensor surface. At present, SiPM sensors are one of the most expensive components of a PET ring.
Methods that measure DOI information are currently not yet implemented with SiPMs on only one crystal side of a monolithic crystal. The concept that is implemented in [5, 6] uses position-sensitive PMTs that cannot be used in strong magnetic fields. They are therefore not MRT-compatible. The concept could be implemented with MRT-compatible avalanche photodiodes (APDs), which has not yet been done to date. APDs are photodiodes that undergo an avalanche effect in which photoelectrons generated by light are accelerated and activate more electrons. The resulting photocurrent depends on the light intensity, as is the case with PMTs. Nevertheless, the implementation of this concept at the level of SiPM microcells poses another challenge, since SiPM microcells are binary sensors and are operated in a different mode, the so-called Geiger mode.
A three-dimensional animal PET scanner was integrated by Judenhofer et al. [8] into a 7 T animal scanner. It is based on APDs that use scintillation crystals with a thickness of 4.5 mm and consisting of crystal arrays with 144 crystals having a pitch of 1.6 mm. The crystal array is coupled with a 3×3-sized APD array. The axial field of view (FOV) is 19 mm. This developed system shows that space is very limited, especially for integrated systems, which forces a compromise between crystal thickness and axial FOV. This results in the system's low sensitivity of 0.23%. Moreover, the DOI problem limits the crystal thickness here, too.
The first version of another prototype scanner that was made public under the name MADPET was developed in Munich [9]. It is made with APDs that have been coupled directly with 3.7 mm×3.7 mm×12 mm crystals. This prototype scanner exhibits the problem of an increase in readout channels when using 1-to-1 coupling. In the first scanner, it is not possible to read all channels at the same time. What is more, the scanner has the problem of low sensitivity. In a second version of the scanner, MADPET II, this problem was solved, and it is possible to read out all of the APDs [14]. The second version also has a two-layered readout system with two layers of crystals with interposed APDs. Since the crystals are therefore partitioned, DOI positions can also be determined. However, twice the amount of sensor surface is also required, which raises the number of readout channels again. Moreover, the approximately doubled quantity of sensors results in higher costs.
The possibility of DOI detection with position-sensitive PMTs was established in [10, 11].
Research findings with detectors consisting of SiPMs and monolithic crystals are published in [12]. In this approach, SiPMs are used in the same manner as the original concept for PMTs and APDs published in [5, 6].